A normal ear transmits sounds as shown in FIG. 1 through the outer ear 101 to the tympanic membrane 102, which moves the bones of the middle ear 103 (malleus, incus, and stapes) that vibrate the oval window and round window openings of the cochlea 104. The cochlea 104 is a long narrow duct wound spirally about its axis for approximately two and a half turns. It includes an upper channel known as the scala vestibuli and a lower channel known as the scala tympani, which are connected by the cochlear duct. The cochlea 104 forms an upright spiraling cone with a center called the modiolar where the spiral ganglion cells of the acoustic nerve 113 reside. In response to received sounds transmitted by the middle ear 103, the fluid-filled cochlea 104 functions as a transducer to generate electric pulses which are transmitted to the cochlear nerve 113, and ultimately to the brain.
Hearing is impaired when there are problems in the ability to transduce external sounds into meaningful action potentials along the neural substrate of the cochlea 104. To improve impaired hearing, hearing prostheses have been developed. For example, when the impairment is related to operation of the middle ear 103, a conventional hearing aid may be used to provide mechanical stimulation to the auditory system in the form of amplified sound. Or when the impairment is associated with the cochlea 104, a cochlear implant with an implanted stimulation electrode can electrically stimulate auditory nerve tissue with small currents delivered by multiple electrode contacts distributed along the electrode.
In some patients with some residual hearing in the lower acoustic frequencies, a conventional hearing aid and a cochlear implant can be combined together in a hybrid Electric Acoustic Stimulation (EAS) system. The hearing aid acoustically amplifies lower acoustic frequencies perceived by human ear, while the cochlear implant electrically stimulates the middle and high frequencies. See von Ilberg et al, Electric-Acoustic Stimulation of the Auditory System, ORL 61:334-340; Skarzynski et al, Preservation of Low Frequency Hearing in Partial Deafness Cochlear Implantation (PDCI) Using the Round Window Surgical Approach, Acta OtoLaryngol 2007; 127:41-48; Gantz & Turner, Combining Acoustic and Electrical Speech Processing: Iowa/Nucleus Hybrid Implant, Acta Otolaryngol 2004; 124:344-347; Gstottner et al., Hearing Preservation in Cochlear Implantation for Electric Acoustic Stimulation, Acta Otolaryngol 2004; 124:348-352; all incorporated herein by reference.
FIG. 1 also shows some components of a typical EAS system which includes an external microphone that provides an acoustic signal input to an external signal processor 111 where two different signal processing paths are developed. An upper acoustic frequency range communications signal containing middle and high frequency range acoustic is converted into a digital data format, such as a sequence of data frames, for transmission via a transmitter coil 107 over a corresponding implanted receiver coil 106 into the electric implant 108 (a typical cochlear implant system). Besides receiving the processed acoustic information, the electric implant 108 also performs additional signal processing such as error correction, pulse formation, etc., and produces an electric stimulation pattern (based on the extracted acoustic information) that is sent through an electrode lead 109 to an implanted electrode array 110. Typically, this electrode array 110 includes multiple electrode contacts on its outer surface that provide selective electric stimulation of the cochlea 104. The external signal processor 111 also creates a lower acoustic frequency range communications signal to a conventional hearing aid 105 in the ear canal which acoustically stimulates the tympanic membrane 102, and in turn the middle ear 103 and cochlea 104.
In some coding strategies, stimulation pulses are applied at a constant rate across all electrode channels, whereas in other coding strategies, stimulation pulses are applied at a channel-specific rate. Various specific signal processing schemes can be implemented to produce the electrical stimulation signals. Signal processing approaches that are well-known in the field of cochlear implants include continuous interleaved sampling (CIS), channel specific sampling sequences (CSSS) (as described in U.S. Pat. No. 6,348,070, incorporated herein by reference), spectral peak (SPEAK), and compressed analog (CA) processing.
FIG. 2 shows the major functional blocks in a typical cochlear implant signal processing system wherein band pass signals are processed and coding to generate electrode stimulation signals to stimulation electrodes in an implanted cochlear implant electrode array. For example, commercially available Digital Signal Processors (DSP) can be used to perform speech processing according to a 12-channel CIS approach. The initial acoustic audio signal input is produced by one or more sensing microphones, which may be omnidirectional and/or directional. Preprocessor Filter Bank 201 pre-processes the initial acoustic audio signal with a bank of multiple band pass filters, each of which is associated with a specific band of audio frequencies—for example, a digital filter bank having 12 digital Butterworth band pass filters of 6th order, Infinite Impulse Response (IIR) type—so that the acoustic audio signal is filtered into some N band pass signals, B1 to BN where each signal corresponds to the band of frequencies for one of the band pass filters. Each output of the CIS band pass filters can roughly be regarded as a sinusoid at the center frequency of the band pass filter which is modulated by the envelope signal. This is due to the quality factor (Q≈3) of the filters. In case of a voiced speech segment, this envelope is approximately periodic, and the repetition rate is equal to the pitch frequency. Alternatively and without limitation, the Preprocessor Filter Bank 201 may be implemented based on use of a fast Fourier transform (FFT) or a short-time Fourier transform (STFT). Based on the tonotopic organization of the cochlea, each electrode contact in the scala tympani often is associated with a specific band pass filter of the external filter bank.
FIG. 3 shows an example of a short time period of an audio speech signal from a microphone, and FIG. 4 shows an acoustic microphone signal decomposed by band-pass filtering by a bank of filters into a set of signals. An example of pseudocode for an infinite impulse response (IIR) filter bank based on a direct form II transposed structure is given by Fontaine et al., Brian Hears: Online Auditory Processing Using Vectorization Over Channels, Frontiers in Neuroinformatics, 2011; incorporated herein by reference in its entirety:
for j = 0 to number of channels - 1 do for s = 0 to number of samples - 1 do  Yj(s) = B0j * Xj (s) + Z0j  for i = 0 to order - 3 do   Zij = Bi+1,j * Xj(s) + Zi+1,j - Ai+1,j * Yj (s)  end for  Zorder-2,j = Border-1,j * Xj(s) - Aorder-1,j * Yj (s) end forend for
The band pass signals B1 to BN (which can also be thought of as frequency channels) are input to a Signal Processor 202 which extracts signal specific stimulation information—e.g., envelope information, phase information, timing of requested stimulation events, etc.—into a set of N stimulation channel signals S1 to SN that represent electrode specific requested stimulation events. For example, channel specific sampling sequences (CSSS) may be used as described in U.S. Pat. No. 6,594,525, which is incorporated herein by reference in its entirety. For example, the envelope extraction may be performed using 12 rectifiers and 12 digital Butterworth low pass filters of 2nd order, IIR-type.
A Pulse Generator 205 includes a Pulse Mapping Module 203 that applies a non-linear mapping function (typically logarithmic) to the amplitude of each band-pass envelope. This mapping function—for example, using instantaneous nonlinear compression of the envelope signal (map law)—typically is adapted to the needs of the individual cochlear implant user during fitting of the implant in order to achieve natural loudness growth. This may be in the specific form of functions that are applied to each requested stimulation event signal S1 to SN that reflect patient-specific perceptual characteristics to produce a set of electrode stimulation signals A1 to AM that provide an optimal electric representation of the acoustic signal. A logarithmic function with a form-factor C typically may be applied as a loudness mapping function, which typically is identical across all the band pass analysis channels. In different systems, different specific loudness mapping functions other than a logarithmic function may be used, with just one identical function is applied to all channels or one individual function for each channel to produce the electrode stimulation signals A1 to AM outputs from the Pulse Mapping Module 203.
The Pulse Generator 205 also includes a Pulse Shaper 204 that develops the set of electrode stimulation signals A1 to AM into a set of output electrode pulses E1 to EM for the electrode contacts in the implanted electrode array which stimulate the adjacent nerve tissue. The electrode stimulation signals A1 to AM may be symmetrical biphasic current pulses with amplitudes that are directly obtained from the compressed envelope signals.
In the specific case of a CIS system, the stimulation pulses are applied in a strictly non-overlapping sequence. Thus, as a typical CIS-feature, only one electrode channel is active at a time and the overall stimulation rate is comparatively high. For example, assuming an overall stimulation rate of 18 kpps and a 12 channel filter bank, the stimulation rate per channel is 1.5 kpps. Such a stimulation rate per channel usually is sufficient for adequate temporal representation of the envelope signal. The maximum overall stimulation rate is limited by the minimum phase duration per pulse. The phase duration cannot be arbitrarily short because, the shorter the pulses, the higher the current amplitudes have to be to elicit action potentials in neurons, and current amplitudes are limited for various practical reasons. For an overall stimulation rate of 18 kpps, the phase duration is 27 μs, which is near the lower limit.
In the CIS strategy, the signal processor only uses the band pass signal envelopes for further processing, i.e., they contain the entire stimulation information. For each electrode channel, the signal envelope is represented as a sequence of biphasic pulses at a constant repetition rate. A characteristic feature of CIS is that the stimulation rate is equal for all electrode channels and there is no relation to the center frequencies of the individual channels. It is intended that the pulse repetition rate is not a temporal cue for the patient (i.e., it should be sufficiently high so that the patient does not perceive tones with a frequency equal to the pulse repetition rate). The pulse repetition rate is usually chosen at greater than twice the bandwidth of the envelope signals (based on the Nyquist theorem).
Another cochlear implant stimulation strategy that does transmit fine time structure information is the Fine Structure Processing (FSP) strategy by Med-El. Zero crossings of the band pass filtered time signals are tracked, and at each negative to positive zero crossing, a Channel Specific Sampling Sequence (CSSS) is started. Typically CSSS sequences are only applied on the first one or two most apical electrode channels, covering the frequency range up to 200 or 330 Hz. The FSP arrangement is described further in Hochmair I, Nopp P, Jolly C, Schmidt M, Schöβer H, Garnham C, Anderson I, MED-EL Cochlear Implants: State of the Art and a Glimpse into the Future, Trends in Amplification, vol. 10, 201-219, 2006, which is incorporated herein by reference.
Many cochlear implant coding strategies use what is referred to as an N-of-M approach where only some number n electrode channels with the greatest amplitude are stimulated in a given sampling time frame. If, for a given time frame, the amplitude of a specific electrode channel remains higher than the amplitudes of other channels, then that channel will be selected for the whole time frame. Subsequently, the number of electrode channels that are available for coding information is reduced by one, which results in a clustering of stimulation pulses. Thus, fewer electrode channels are available for coding important temporal and spectral properties of the sound signal such as speech onset.
One method to reduce the spectral clustering of stimulation per time frame is the MP3000™ coding strategy by Cochlear Ltd, which uses a spectral masking model on the electrode channels. Another method that inherently enhances coding of speech onsets is the ClearVoice™ coding strategy used by Advanced Bionics Corp, which selects electrode channels having a high signal to noise ratio. U.S. Patent Publication 2005/0203589 (which is incorporated herein by reference in its entirety) describes how to organize electrode channels into two or more groups per time frame. The decision which electrode channels to select is based on the amplitude of the signal envelopes.
In addition to the specific processing and coding approaches discussed above, different specific pulse stimulation modes are possible to deliver the stimulation pulses with specific electrodes—i.e. mono-polar, bi-polar, tri-polar, multi-polar, and phased-array stimulation. And there also are different stimulation pulse shapes—i.e. biphasic, symmetric triphasic, asymmetric triphasic pulses, or asymmetric pulse shapes. These various pulse stimulation modes and pulse shapes each provide different benefits; for example, higher tonotopic selectivity, smaller electrical thresholds, higher electric dynamic range, less unwanted side-effects such as facial nerve stimulation, etc. But some stimulation arrangements are quite power consuming, especially when neighboring electrodes are used as current sinks. Up to 10 dB more charge might be required than with simple mono-polar stimulation concepts (if the power-consuming pulse shapes or stimulation modes are used continuously).
It is well-known in the field that electric stimulation at different locations within the cochlea produce different frequency percepts. The underlying mechanism in normal acoustic hearing is referred to as the tonotopic principle. In cochlear implant users, the tonotopic organization of the cochlea has been extensively investigated; for example, see Vermeire et al., Neural tonotopy in cochlear implants: An evaluation in unilateral cochlear implant patients with unilateral deafness and tinnitus, Hear Res, 245(1-2), 2008 Sep. 12 p. 98-106; and Schatzer et al., Electric-acoustic pitch comparisons in single-sided-deaf cochlear implant users: Frequency-place functions and rate pitch, Hear Res, 309, 2014 March, p. 26-35 (both of which are incorporated herein by reference in their entireties). According to the Greenwood scale, 360° of electrode insertion in the tonotopically organized cochlea covers the acoustic frequency region from 1 kHz and higher. See Greenwood, A Cochlear Frequency-Position Function For Several Species 29 Years Later, J Acoustic Soc Am, 1990; 87(6):2592-2605; incorporated herein by reference.
In a normal hearing ear, one frequency component consecutively stimulates multiple neural populations. This phenomenon was described as the “travelling wave” as shown in FIG. 5 from Von Békésy, Georg. Experiments in hearing. Ed. Ernest Glen Wever. Vol. 8. New York: McGraw-Hill, 1960 (incorporated herein by reference in its entirety). That is, in response to a pure tone, the basilar membrane resonates in a travelling wave (the ascending numbers within FIG. 5) which gradually grows in amplitude (the dashed lines in FIG. 5) as it moves along the cochlear duct from the stapes (base) toward the helicotrema (apex).
One quality of the travelling wave that is partly reflected in modern cochlear implant systems is that each frequency component reaches a peak amplitude at a specific spot within the cochlea (the tonotopic principle discussed above). These spectro-temporal properties can also be observed in the activity of cat's cochlear nerve fibres shown in FIG. 6 from Secker Walker et al, Time domain analysis of auditory nerve fiber firing rates, J Acoust Soc Am, 88(3), 1990, p. 1427-1436 (incorporated herein by reference in its entirety). FIG. 6 shows neural activity in the cochlear nerve over time at nerve fibres with different characteristic frequencies in response to synthetic vowels. One dominant frequency component in the synthetic vowel stimuli is the fundamental frequency (F0), which in FIG. 6 can be clearly identified as a regular pattern starting at high frequencies and ending several milliseconds later at low frequencies. The black curve in the shaded box in FIG. 6 indicates the frequency-specific time delays or the neural responses. Higher frequency components also can be observed between the F0 structures; for example, harmonics that are visible between 1800 and 1000 Hz. Similar to the F0 structures, they start at high frequency fibers and end some milliseconds later at low frequency fibers. This spectro-temporal excitation behaviour is not currently explicitly implemented in cochlear implant systems.
Loeb G., Are cochlear implant patients suffering from perceptual dissonance? Ear Hear, 26, 2005, p. 435-450 (incorporated herein by reference in its entirety) describes that phase-locking occurs over a substantial length of the cochlea. Furthermore, the action potentials exhibit a coherent spatial gradient with the steepest and most rapidly changing gradient of the phase occurring next to the place of the resonant frequency. At this point, the travelling wave starts to significantly slow down and dissipates. The phase gradient is believed to substantially contribute to pitch perception, especially in loud situations where harmonics are not resolved.
Existing coding approaches take into account some of the temporal properties of the acoustic signal. CIS determines frequency-specific envelopes which inherently contain a certain amount of information about individual low frequency components such as the fundamental frequency. More advanced approaches for calculating band specific envelopes also have been described; for example, U.S. Patent Publication 2006/0235486 (which is incorporated herein by reference in its entirety). The latter and CIS both sample the band pass envelopes with fixed rate stimulation pulses to resemble rudimentary properties of the basilar membrane movement. Other advanced systems as described in U.S. Patent Publication 2011/0230934 (which is incorporated herein by reference in its entirety) explicitly extract temporal characteristics of a band pass signal by identifying phase characteristics such as zero crossings. The described system triggers channel-specific sequences of stimulation pulses at each detected zero crossing. Each of the foregoing arrangements attributes certain frequency components to certain stimulation places. U.S. Patent Publication 2011/0230934 also explicitly takes into account the timing of certain frequency components.
Vocoder-based cochlear implant stimulation arrangements such as CIS and N-of-M do not take into account the travelling wave properties of normal acoustic hearing. The acoustic signal is analysed by filter banks or FFT and assigned either to single intracochlear electrodes, or to simultaneous stimulation of multiple adjacent electrodes. While filter banks can mimic the latencies of single frequency components at the place of stimulation, they are not able to mimic other aspects of the travelling wave behaviour such as the spectro-temporal distribution of this component to neighbouring stimulation sites, starting at a more basal site with low amplitude and ending at a more apical stimulation site with a maximum of stimulation at a site in between. An FFT, also used for mimicking the tonotopic principle in a cochlear implant is no better able to replicate the general latency differences between the frequency components (at the place of stimulation) nor does it provide the spectro-temporal behaviour described above.
In those patients receiving hearing implants who retain significant post-surgical residual natural hearing, typically, the residual hearing is in the lower frequency bands, which tonotopically corresponds to the deepest locations within the cochlea. In some cases, an apical portion of the implanted electrode array reaches far enough in to the cochlea to enter into the regions with residual hearing. In such cases, it has been suggested that the most apical electrode channels be turned off. For example, U.S. 2013/0116746 suggests that the electrode array be inserted only to a shallow depth in the scala tympani which typically is a region without residual hearing. Or if the most apical contacts of the electrode array actually are located in the acoustically perceivable region of the patient, to intentionally introduce a frequency gap between the acoustically perceivable frequencies and the electrically perceivable frequencies by turning off one or more of the most apical electrode contacts to create a (small) region without any hearing stimulation, either acoustic or electric.
U.S. Pat. No. 8,000,798 suggests an electrode array that after insertion has a subset of electrode contacts beyond a first basal turn of the cochlea, and activating one or more of the contacts in this subset while simultaneously allowing natural acoustic hearing to occur in one or more locations beyond the first basal turn. But it is apparent from the description that this patent describes “a relatively very thin and short electrode array that is insertable into the basal region of the cochlea and past the first turn thereof”
WO 00/69513 suggests a hybrid cochlear stimulation arrangement (i.e., combined electrical stimulation together with acoustic stimulation of residual hearing) that uses an electrode array which is no longer than 8 mm—shorter than the distance from the cochleostomy or round window membrane to the first basal turn.
All these existing measures are based on current beliefs that electrically stimulated hearing is different from natural acoustic hearing, and that electric stimulation interacts with acoustic hearing.